Charged particle cancer therapy system magnet control method and apparatus

ABSTRACT

The invention comprises a charged particle beam acceleration, extraction, and/or targeting method and apparatus used in conjunction with charged particle beam radiation therapy of cancerous tumors. Novel design features of a synchrotron are described. Particularly, turning magnets, edge focusing magnets, concentrating magnetic field magnets, winding and control coils, and extraction elements are described that minimize the overall size of the synchrotron, provide a tightly controlled proton beam, directly reduce the size of required magnetic fields, directly reduces required operating power, and allow continual acceleration of protons in a synchrotron even during a process of extracting protons from the synchrotron.

CROSS REFERENCES TO RELATED APPLICATIONS

This application:

-   -   is a continuation-in-part of U.S. patent application Ser. No.         12/425,683 filed Apr. 17, 2009 now U.S. Pat. No. 7,939,809,         which claims the benefit of:         -   U.S. provisional patent application No. 61/055,395 filed May             22, 2008;         -   U.S. provisional patent application No. 61/137,574 filed             Aug. 1, 2008;         -   U.S. provisional patent application No. 61/192,245 filed             Sep. 17, 2008;         -   U.S. provisional patent application No. 61/055,409 filed May             22, 2008;         -   U.S. provisional patent application No. 61/203,308 filed             Dec. 22, 2008;         -   U.S. provisional patent application No. 61/188,407 filed             Aug. 11, 2008;         -   U.S. provisional patent application No. 61/188,406 filed             Aug. 11, 2008;         -   U.S. provisional patent application No. 61/189,815 filed             Aug. 25, 2008;         -   U.S. provisional patent application No. 61/201,731 filed             Dec. 15, 2008;         -   U.S. provisional patent application No. 61/205,362 filed             Jan. 12, 2009;         -   U.S. provisional patent application No. 61/134,717 filed             Jul. 14, 2008;         -   U.S. provisional patent application No. 61/134,707 filed             Jul. 14, 2008;         -   U.S. provisional patent application No. 61/201,732 filed             Dec. 15, 2008;         -   U.S. provisional patent application No. 61/198,509 filed             Nov. 7, 2008;         -   U.S. provisional patent application No. 61/134,718 filed             Jul. 14, 2008;         -   U.S. provisional patent application No. 61/190,613 filed             Sep. 2, 2008;         -   U.S. provisional patent application No. 61/191,043 filed             Sep. 8, 2008;         -   U.S. provisional patent application No. 61/192,237 filed             Sep. 17, 2008;         -   U.S. provisional patent application No. 61/201,728 filed             Dec. 15, 2008;         -   U.S. provisional patent application No. 61/190,546 filed             Sep. 2, 2008;         -   U.S. provisional patent application No. 61/189,017 filed             Aug. 15, 2008;         -   U.S. provisional patent application No. 61/198,248 filed             Nov. 5, 2008;         -   U.S. provisional patent application No. 61/198,508 filed             Nov. 7, 2008;         -   U.S. provisional patent application No. 61/197,971 filed             Nov. 3, 2008;         -   U.S. provisional patent application No. 61/199,405 filed             Nov. 17, 2008;         -   U.S. provisional patent application No. 61/199,403 filed             Nov. 17, 2008; and         -   U.S. provisional patent application No. 61/199,404 filed             Nov. 17, 2008;     -   claims the benefit of U.S. provisional patent application No.         61/209,529 filed Mar. 9, 2009;     -   claims the benefit of U.S. provisional patent application No.         61/208,182 filed Feb. 23, 2009;     -   claims the benefit of U.S. provisional patent application No.         61/208,971 filed Mar. 3, 2009; and     -   claims priority to PCT patent application serial No.:         PCT/RU2009/00015, filed Mar. 4, 2009,     -   all of which are incorporated herein in their entirety by this         reference thereto.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates generally to treatment of solid cancers. More particularly, the invention relates to a charged particle beam acceleration, extraction, and/or targeting method and apparatus used in conjunction with radiation treatment of cancerous tumors.

2. Discussion of the Prior Art

Cancer

A tumor is an abnormal mass of tissue. Tumors are either benign or malignant. A benign tumor grows locally, but does not spread to other parts of the body. Benign tumors cause problems because of their spread, as they press and displace normal tissues. Benign tumors are dangerous in confined places such as the skull. A malignant tumor is capable of invading other regions of the body. Metastasis is cancer spreading by invading normal tissue and spreading to distant tissues.

Cancer Treatment

Several forms of radiation therapy exist for cancer treatment including: brachytherapy, traditional electromagnetic X-ray therapy, and proton therapy. Each are further described, infra.

Brachytherapy is radiation therapy using radioactive sources implanted inside the body. In this treatment, an oncologist implants radioactive material directly into the tumor or very close to it. Radioactive sources are also placed within body cavities, such as the uterine cervix.

The second form of traditional cancer treatment using electromagnetic radiation includes treatment using X-rays and gamma rays. An X-ray is high-energy, ionizing, electromagnetic radiation that is used at low doses to diagnose disease or at high doses to treat cancer. An X-ray or Röntgen ray is a form of electromagnetic radiation with a wavelength in the range of 10 to 0.01 nanometers (nm), corresponding to frequencies in the range of 30 PHz to 30 EHz. X-rays are longer than gamma rays and shorter than ultraviolet rays. X-rays are primarily used for diagnostic radiography. X-rays are a form of ionizing radiation and as such can be dangerous. Gamma rays are also a form of electromagnetic radiation and are at frequencies produced by sub-atomic particle interactions, such as electron-positron annihilation or radioactive decay. In the electromagnetic spectrum, gamma rays are generally characterized as electromagnetic radiation having the highest frequency, as having highest energy, and having the shortest wavelength, such as below about 10 picometers. Gamma rays consist of high energy photons with energies above about 100 keV. X-rays are commonly used to treat cancerous tumors. However, X-rays are not optimal for treatment of cancerous tissue as X-rays deposit their highest does of radiation near the surface of the targeted tissue and delivery exponentially less radiation as they penetrate into the tissue. This results in large amounts of radiation being delivered outside of the tumor. Gamma rays have similar limitations.

The third form of cancer treatment uses protons. Proton therapy systems typically include: a beam generator, an accelerator, and a beam transport system to move the resulting accelerated protons to a plurality of treatment rooms where the protons are delivered to a tumor in a patient's body.

Proton therapy works by aiming energetic ionizing particles, such as protons accelerated with a particle accelerator, onto a target tumor. These particles damage the DNA of cells, ultimately causing their death. Cancerous cells, because of their high rate of division and their reduced ability to repair damaged DNA, are particularly vulnerable to attack on their DNA.

Due to their relatively enormous size, protons scatter less easily in the tissue and there is very little lateral dispersion. Hence, the proton beam stays focused on the tumor shape without much lateral damage to surrounding tissue. All protons of a given energy have a certain range, defined by the Bragg peak, and the dosage delivery to tissue ratio is maximum over just the last few millimeters of the particle's range. The penetration depth depends on the energy of the particles, which is directly related to the speed to which the particles were accelerated by the proton accelerator. The speed of the proton is adjustable to the maximum rating of the accelerator. It is therefore possible to focus the cell damage due to the proton beam at the very depth in the tissues where the tumor is situated. Tissues situated before the Bragg peak receive some reduced dose and tissues situated after the peak receive none.

Synchrotrons

Patents related to the current invention are summarized here.

Proton Beam Therapy System

F. Cole, et. al. of Loma Linda University Medical Center “Multi-Station Proton Beam Therapy System”, U.S. Pat. No. 4,870,287 (Sep. 26, 1989) describe a proton beam therapy system for selectively generating and transporting proton beams from a single proton source and accelerator to a selected treatment room of a plurality of patient treatment rooms.

Injection

K. Hiramoto, et. al. “Accelerator System”, U.S. Pat. No. 4,870,287 (Sep. 26, 1989) describes an accelerator system having a selector electromagnet for introducing an ion beam accelerated by pre-accelerators into either a radioisotope producing unit or a synchrotron.

K. Hiramoto, et. al. “Circular Accelerator, Method of Injection of Charged Particle Thereof, and Apparatus for Injection of Charged Particle Thereof”, U.S. Pat. No. 5,789,875 (Aug. 4, 1998) and K. Hiramoto, et. al. “Circular Accelerator, Method of Injection of Charged Particle Thereof, and Apparatus for Injection of Charged Particle Thereof”, U.S. Pat. No. 5,600,213 (Feb. 4, 1997) both describe a method and apparatus for injecting a large number of charged particles into a vacuum duct where the beam of injection has a height and width relative to a geometrical center of the duct.

Accelerator/Synchrotron

H. Tanaka, et. al. “Charged Particle Accelerator”, U.S. Pat. No. 7,259,529 (Aug. 21, 2007) describe a charged particle accelerator having a two period acceleration process with a fixed magnetic field applied in the first period and a timed second acceleration period to provide compact and high power acceleration of the charged particles.

T. Haberer, et. al. “Ion Beam Therapy System and a Method for Operating the System”, U.S. Pat. No. 6,683,318 (Jan. 27, 2004) describe an ion beam therapy system and method for operating the system. The ion beam system uses a gantry that has vertical deflection system and a horizontal deflection system positioned before a last bending magnet that result in a parallel scanning mode resulting from an edge focusing effect.

V. Kulish, et. al. “Inductional Undulative EH-Accelerator”, U.S. Pat. No. 6,433,494 (Aug. 13, 2002) describe an inductive undulative EH-accelerator for acceleration of beams of charged particles. The device consists of an electromagnet undulation system, whose driving system for electromagnets is made in the form of a radio-frequency (RF) oscillator operating in the frequency range from about 100 KHz to 10 GHz.

K. Saito, et. al. “Radio-Frequency Accelerating System and Ring Type Accelerator Provided with the Same”, U.S. Pat. No. 5,917,293 (Jun. 29, 1999) describe a radio-frequency accelerating system having a loop antenna coupled to a magnetic core group and impedance adjusting means connected to the loop antenna. A relatively low voltage is applied to the impedance adjusting means allowing small construction of the adjusting means.

J. Hirota, et. al. “Ion Beam Accelerating Device Having Separately Excited Magnetic Cores”, U.S. Pat. No. 5,661,366 (Aug. 26, 1997) describe an ion beam accelerating device having a plurality of high frequency magnetic field inducing units and magnetic cores.

J. Hirota, et. al. “Acceleration Device for Charged Particles”, U.S. Pat. No. 5,168,241 (Dec. 1, 1992) describe an acceleration cavity having a high frequency power source and a looped conductor operating under a control that combine to control a coupling constant and/or de-tuning allowing transmission of power more efficiently to the particles.

Vacuum Chamber

T. Kobari, et. al. “Apparatus For Treating the Inner Surface of Vacuum Chamber”, U.S. Pat. No. 5,820,320 (Oct. 13, 1998) and T. Kobari, et. al. “Process and Apparatus for Treating Inner Surface Treatment of Chamber and Vacuum Chamber”, U.S. Pat. No. 5,626,682 (May 6, 1997) both describe an apparatus for treating an inner surface of a vacuum chamber including means for supplying an inert gas or nitrogen to a surface of the vacuum chamber with a broach. Alternatively, the broach is used for supplying a lower alcohol to the vacuum chamber for dissolving contaminants on the surface of the vacuum chamber.

Magnet Shape

M. Tadokoro, et. al. “Electromagnetic and Magnetic Field Generating Apparatus”, U.S. Pat. No. 6,365,894 (Apr. 2, 2002) and M. Tadokoro, et. al. “Electromagnetic and Magnetic Field Generating Apparatus”, U.S. Pat. No. 6,236,043 (May 22, 2001) each describe a pair of magnetic poles, a return yoke, and exciting coils. The interior of the magnetic poles each have a plurality of air gap spacers to increase magnetic field strength.

Extraction

T. Nakanishi, et. al. “Charged-Particle Beam Accelerator, Particle Beam Radiation Therapy System Using the Charged-Particle Beam Accelerator, and Method of Operating the Particle Beam Radiation Therapy System”, U.S. Pat. No. 7,122,978 (Oct. 17, 2006) describe a charged particle beam accelerator having an RF-KO unit for increasing amplitude of betatron oscillation of a charged particle beam within a stable region of resonance and an extraction quadrupole electromagnet unit for varying a stable region of resonance. The RF-KO unit is operated within a frequency range in which the circulating beam does not go beyond a boundary of stable region of resonance and the extraction quadrupole electromagnet is operated with timing required for beam extraction.

T. Haberer, et. al. “Method and Device for Controlling a Beam Extraction Raster Scan Irradiation Device for Heavy Ions or Protons”, U.S. Pat. No. 7,091,478 (Aug. 15, 2006) describe a method for controlling beam extraction irradiation in terms of beam energy, beam focusing, and beam intensity for every accelerator cycle.

K. Hiramoto, et. al. “Accelerator and Medical System and Operating Method of the Same”, U.S. Pat. No. 6,472,834 (Oct. 29, 2002) describe a cyclic type accelerator having a deflection electromagnet and four-pole electromagnets for making a charged particle beam circulate, a multi-pole electromagnet for generating a stability limit of resonance of betatron oscillation, and a high frequency source for applying a high frequency electromagnetic field to the beam to move the beam to the outside of the stability limit. The high frequency source generates a sum signal of a plurality of alternating current (AC) signals of which the instantaneous frequencies change with respect to time, and of which the average values of the instantaneous frequencies with respect to time are different. The system applies the sum signal via electrodes to the beam.

K. Hiramoto, et. al. “Synchrotron Type Accelerator and Medical Treatment System Employing the Same”, U.S. Pat. No. 6,087,670 (Jul. 11, 2000) and K. Hiramoto, et. al. “Synchrotron Type Accelerator and Medical Treatment System Employing the Same”, U.S. Pat. No. 6,008,499 (Dec. 28, 1999) describe a synchrotron accelerator having a high frequency applying unit arranged on a circulating orbit for applying a high frequency electromagnetic field to a charged particle beam circulating and for increasing amplitude of betatron oscillation of the particle beam to a level above a stability limit of resonance. Additionally, for beam ejection, four-pole divergence electromagnets are arranged: (1) downstream with respect to a first deflector; (2) upstream with respect to a deflecting electromagnet; (3) downstream with respect to the deflecting electromagnet; and (4) and upstream with respect to a second deflector.

K. Hiramoto, et. al. “Circular Accelerator and Method and Apparatus for Extracting Charged-Particle Beam in Circular Accelerator”, U.S. Pat. No. 5,363,008 (Nov. 8, 1994) describe a circular accelerator for extracting a charged-particle beam that is arranged to: (1) increase displacement of a beam by the effect of betatron oscillation resonance; (2) to increase the betatron oscillation amplitude of the particles, which have an initial betatron oscillation within a stability limit for resonance; and (3) to exceed the resonance stability limit thereby extracting the particles exceeding the stability limit of the resonance.

K. Hiramoto, et. al. “Method of Extracting Charged Particles from Accelerator, and Accelerator Capable Carrying Out the Method, by Shifting Particle Orbit”, U.S. Pat. No. 5,285,166 (Feb. 8, 1994) describe a method of extracting a charged particle beam. An equilibrium orbit of charged particles maintained by a bending magnet and magnets having multipole components greater than sextuple components is shifted by a constituent element of the accelerator other than these magnets to change the tune of the charged particles.

Transport/Scanning Control

K. Matsuda, et. al. “Particle Beam Irradiation Apparatus, Treatment Planning Unit, and Particle Beam Irradiation Method”, U.S. Pat. No. 7,227,161 (Jun. 5, 2007); K. Matsuda, et. al. “Particle Beam Irradiation Treatment Planning Unit, and Particle Beam Irradiation Method”, U.S. Pat. No. 7,122,811 (Oct. 17, 2006); and K. Matsuda, et. al. “Particle Beam Irradiation Apparatus, Treatment Planning Unit, and Particle Beam Irradiation Method” (Sep. 5, 2006) describe a particle beam irradiation apparatus have a scanning controller that stops output of an ion beam, changes irradiation position via control of scanning electromagnets, and reinitiates treatment based on treatment planning information.

T. Norimine, et. al. “Particle Therapy System Apparatus”, U.S. Pat. Nos. 7,060,997 (Jun. 13, 2006); T. Norimine, et. al. “Particle Therapy System Apparatus”, 6,936,832 (Aug. 30, 2005); and T. Norimine, et. al. “Particle Therapy System Apparatus”, 6,774,383 (Aug. 10, 2004) each describe a particle therapy system having a first steering magnet and a second steering magnet disposed in a charged particle beam path after a synchrotron that are controlled by first and second beam position monitors.

K. Moriyama, et. al. “Particle Beam Therapy System”, U.S. Pat. No. 7,012,267 (Mar. 14, 2006) describe a manual input to a ready signal indicating preparations are completed for transport of the ion beam to a patient.

H. Harada, et. al. “Irradiation Apparatus and Irradiation Method”, U.S. Pat. No. 6,984,835 (Jan. 10, 2006) describe an irradiation method having a large irradiation filed capable of uniform dose distribution, without strengthening performance of an irradiation field device, using a position controller having overlapping area formed by a plurality of irradiations using a multileaf collimator. The system provides flat and uniform dose distribution over an entire surface of a target.

H. Akiyama, et. al. “Charged Particle Beam Irradiation Equipment Having Scanning Electromagnet Power Supplies”, U.S. Pat. No. 6,903,351 (Jun. 7, 2005); H. Akiyama, et. al. “Charged Particle Beam Irradiation Equipment Having Scanning Electromagnet Power Supplies”, U.S. Pat. No. 6,900,436 (May 31, 2005); and H. Akiyama, et. al. “Charged Particle Beam Irradiation Equipment Having Scanning Electromagnet Power Supplies”, U.S. Pat. No. 6,881,970 (Apr. 19, 2005) all describe a power supply for applying a voltage to a scanning electromagnet for deflecting a charged particle beam and a second power supply without a pulsating component to control the scanning electromagnet more precisely allowing for uniform irradiation of the irradiation object.

K. Amemiya, et. al. “Accelerator System and Medical Accelerator Facility”, U.S. Pat. No. 6,800,866 (Oct. 5, 2004) describe an accelerator system having a wide ion beam control current range capable of operating with low power consumption and having a long maintenance interval.

A. Dolinskii, et. al. “Gantry with an Ion-Optical System”, U.S. Pat. No. 6,476,403 (Nov. 5, 2002) describe a gantry for an ion-optical system comprising an ion source and three bending magnets for deflecting an ion beam about an axis of rotation. A plurality of quadrupoles are also provided along the beam path to create a fully achromatic beam transport and an ion beam with difference emittances in the horizontal and vertical planes. Further, two scanning magnets are provided between the second and third bending magnets to direct the beam.

H. Akiyama, et. al. “Charged Particle Beam Irradiation Apparatus”, U.S. Pat. No. 6,218,675 (Apr. 17, 2001) describe a charged particle beam irradiation apparatus for irradiating a target with a charged particle beam that include a plurality of scanning electromagnets and a quadrupole electromagnet between two of the plurality of scanning electromagnets.

K. Matsuda, et. al. “Charged Particle Beam Irradiation System and Method Thereof”, U.S. Pat. No. 6,087,672 (Jul. 11, 2000) describe a charged particle beam irradiation system having a ridge filter with shielding elements to shield a part of the charged particle beam in an area corresponding to a thin region in said target.

P. Young, et. al. “Raster Scan Control System for a Charged-Particle Beam”, U.S. Pat. No. 5,017,789 (May 21, 1991) describe a raster scan control system for use with a charged-particle beam delivery system that includes a nozzle through which a charged particle beam passes. The nozzle includes a programmable raster generator and both fast and slow sweep scan electromagnets that cooperate to generate a sweeping magnetic field that steers the beam along a desired raster scan pattern at a target.

Beam Shape Control

M. Yanagisawa, et. al. “Particle Beam Irradiation System and Method of Adjusting Irradiation Field Forming Apparatus”, U.S. Pat. No. 7,154,107 (Dec. 26, 2006) and M. Yanagisawa, et. al. “Particle Beam Irradiation System and Method of Adjusting Irradiation Field Forming Apparatus”, U.S. Pat. No. 7,049,613 (May 23, 2006) describe a particle therapy system having a scattering compensator and a range modulation wheel. Movement of the scattering compensator and the range modulation wheel adjusts a size of the ion beam and scattering intensity resulting in penumbra control and a more uniform dose distribution to a diseased body part.

T. Haberer, et. al. “Device and Method for Adapting the Size of an Ion Beam Spot in the Domain of Tumor Irradiation”, U.S. Pat. No. 6,859,741 (Feb. 22, 2005) describe a method and apparatus for adapting the size of an ion beam in tumor irradiation. Quadrupole magnets determining the size of the ion beam spot are arranged directly in front of raster scanning magnets determining the size of the ion beam spot. The apparatus contains a control loop for obtaining current correction values to further control the ion beam spot size.

K. Matsuda, et. al. “Charged Particle Irradiation Apparatus and an Operating Method Thereof”, U.S. Pat. No. 5,986,274 (Nov. 16, 1999) describe a charged particle irradiation apparatus capable of decreasing a lateral dose falloff at boundaries of an irradiation field of a charged particle beam using controlling magnet fields of quadrupole electromagnets and deflection electromagnets to control the center of the charged particle beam passing through the center of a scatterer irrespective of direction and intensity of a magnetic field generated by scanning electromagnets.

K. Hiramoto, et. al. “Charged Particle Beam Apparatus and Method for Operating the Same”, U.S. Pat. No. 5,969,367 (Oct. 19, 1999) describe a charged particle beam apparatus where a the charged particle beam is enlarged by a scatterer resulting in a Gaussian distribution that allows overlapping of irradiation doses applied to varying spot positions.

M. Moyers, et. al. “Charged Particle Beam Scattering System”, U.S. Pat. No. 5,440,133 (Aug. 8, 1995) describe a radiation treatment apparatus for producing a particle beam and a scattering foil for changing the diameter of the charged particle beam.

C. Nunan “Multileaf Collimator for Radiotherapy Machines”, U.S. Pat. No. 4,868,844 (Sep. 19, 1989) describes a radiation therapy machine having a multileaf collimator formed of a plurality of heavy metal leaf bars movable to form a rectangular irradiation field.

R. Maughan, et. al. “Variable Radiation Collimator”, U.S. Pat. No. 4,754,147 (Jun. 28, 1988) describe a variable collimator for shaping a cross-section of a radiation beam that relies on rods, which are positioned around a beam axis. The rods are shaped by a shaping member cut to a shape of an area of a patient go be irradiated.

Beam Energy/Intensity

M. Yanagisawa, et. al. “Charged Particle Therapy System, Range Modulation Wheel Device, and Method of Installing Range Modulation Wheel Device”, U.S. Pat. No. 7,355,189 (Apr. 8, 2008) and Yanagisawa, et. al. “Charged Particle Therapy System, Range Modulation Wheel Device, and Method of Installing Range Modulation Wheel Device”, U.S. Pat. No. 7,053,389 (May 30, 2008) both describe a particle therapy system having a range modulation wheel. The ion beam passes through the range modulation wheel resulting in a plurality of energy levels corresponding to a plurality of stepped thicknesses of the range modulation wheel.

M. Yanagisawa, et. al. “Particle Beam Irradiation System and Method of Adjusting Irradiation Apparatus”, U.S. Pat. No. 7,297,967 (Nov. 20, 2007); M. Yanagisawa, et. al. “Particle Beam Irradiation System and Method of Adjusting Irradiation Apparatus”, U.S. Pat. No. 7,071,479 (Jul. 4, 2006); M. Yanagisawa, et. al. “Particle Beam Irradiation System and Method of Adjusting Irradiation Apparatus”, U.S. Pat. No. 7,026,636 (Apr. 11, 2006); and M. Yanagisawa, et. al. “Particle Beam Irradiation System and Method of Adjusting Irradiation Apparatus”, U.S. Pat. No. 6,777,700 (Aug. 17, 2004) all describe a scattering device, a range adjustment device, and a peak spreading device. The scattering device and range adjustment device are combined together and are moved along a beam axis. The spreading device is independently moved along the axis to adjust the degree of ion beam scattering. Combined, the devise increases the degree of uniformity of radiation dose distribution to a diseased tissue.

A. Sliski, et. al. “Programmable Particle Scatterer for Radiation Therapy Beam Formation”, U.S. Pat. No. 7,208,748 (Apr. 24, 2007) describe a programmable pathlength of a fluid disposed into a particle beam to modulate scattering angle and beam range in a predetermined manner. The charged particle beam scatterer/range modulator comprises a fluid reservoir having opposing walls in a particle beam path and a drive to adjust the distance between the walls of the fluid reservoir under control of a programmable controller to create a predetermined spread out Bragg peak at a predetermined depth in a tissue. The beam scattering and modulation is continuously and dynamically adjusted during treatment of a tumor to deposit a dose in a targeted predetermined three dimensional volume.

M. Tadokoro, et. al. “Particle Therapy System”, U.S. Pat. No. 7,247,869 (Jul. 24, 2007) and U.S. Pat. No. 7,154,108 (Dec. 26, 2006) each describe a particle therapy system capable of measuring energy of a charged particle beam during irradiation during use. The system includes a beam passage between a pair of collimators, an energy detector mounted, and a signal processing unit.

G. Kraft, et. al. “Ion Beam Scanner System and Operating Method”, U.S. Pat. No. 6,891,177 (May 10, 2005) describe an ion beam scanning system having a mechanical alignment system for the target volume to be scanned and allowing for depth modulation of the ion beam by means of a linear motor and transverse displacement of energy absorption means resulting in depth-staggered scanning of volume elements of a target volume.

G. Hartmann, et. al. “Method for Operating an Ion Beam Therapy System by Monitoring the Distribution of the Radiation Dose”, U.S. Pat. No. 6,736,831 (May 18, 2004) describe a method for operation of an ion beam therapy system having a grid scanner and irradiates and scans an area surrounding an isocentre. Both the depth dose distribution and the transverse dose distribution of the grid scanner device at various positions in the region of the isocentre are measured and evaluated.

Y. Jongen “Method for Treating a Target Volume with a Particle Beam and Device Implementing Same”, U.S. Pat. No. 6,717,162 (Apr. 6, 2004) describes a method of producing from a particle beam a narrow spot directed towards a target volume, characterized in that the spot sweeping speed and particle beam intensity are simultaneously varied.

G. Kraft, et. al. “Device for Irradiating a Tumor Tissue”, U.S. Pat. No. 6,710,362 (Mar. 23, 2004) describe a method and apparatus of irradiating a tumor tissue, where the apparatus has an electromagnetically driven ion-braking device in the proton beam path for depth-wise adaptation of the proton beam that adjusts both the ion beam direction and ion beam range.

K. Matsuda, et. al. “Charged Particle Beam Irradiation Apparatus”, U.S. Pat. No. 6,617,598 (Sep. 9, 2003) describe a charged particle beam irradiation apparatus that increased the width in a depth direction of a Bragg peak by passing the Bragg peak through an enlarging device containing three ion beam components having different energies produced according to the difference between passed positions of each of the filter elements.

H. Stelzer, et. al. “Ionization Chamber for Ion Beams and Method for Monitoring the Intensity of an Ion Beam”, U.S. Pat. No. 6,437,513 (Aug. 20, 2002) describe an ionization chamber for ion beams and a method of monitoring the intensity of an ion therapy beam. The ionization chamber includes a chamber housing, a beam inlet window, a beam outlet window, a beam outlet window, and a chamber volume filled with counting gas.

H. Akiyama, et. al. “Charged-Particle Beam Irradiation Method and System”, U.S. Pat. No. 6,433,349 (Aug. 13, 2002) and H. Akiyama, et. al. “Charged-Particle Beam Irradiation Method and System”, U.S. Pat. No. 6,265,837 (Jul. 24, 2001) both describe a charged particle beam irradiation system that includes a changer for changing energy of the particle and an intensity controller for controlling an intensity of the charged-particle beam.

Y. Pu “Charged Particle Beam Irradiation Apparatus and Method of Irradiation with Charged Particle Beam”, U.S. Pat. No. 6,034,377 (Mar. 7, 2000) describes a charged particle beam irradiation apparatus having an energy degrader comprising: (1) a cylindrical member having a length; and (2) a distribution of wall thickness in a circumferential direction around an axis of rotation, where thickness of the wall determines energy degradation of the irradiation beam.

Dosage

K. Matsuda, et. al. “Particle Beam Irradiation System”, U.S. Pat. No. 7,372,053 (Nov. 27, 2007) describe a particle beam irradiation system ensuring a more uniform dose distribution at an irradiation object through use of a stop signal, which stops the output of the ion beam from the irradiation device.

H. Sakamoto, et. al. “Radiation Treatment Plan Making System and Method”, U.S. Pat. No. 7,054,801 (May 30, 2006) describe a radiation exposure system that divides an exposure region into a plurality of exposure regions and uses a radiation simulation to plan radiation treatment conditions to obtain flat radiation exposure to the desired region.

G. Hartmann, et. al. “Method For Verifying the Calculated Radiation Dose of an Ion Beam Therapy System”, U.S. Pat. No. 6,799,068 (Sep. 28, 2004) describe a method for the verification of the calculated dose of an ion beam therapy system that comprises a phantom and a discrepancy between the calculated radiation dose and the phantom.

H. Brand, et. al. “Method for Monitoring the Irradiation Control of an Ion Beam Therapy System”, U.S. Pat. No. 6,614,038 (Sep. 2, 2003) describe a method of checking a calculated irradiation control unit of an ion beam therapy system, where scan data sets, control computer parameters, measuring sensor parameters, and desired current values of scanner magnets are permanently stored.

T. Kan, et. al. “Water Phantom Type Dose Distribution Determining Apparatus”, U.S. Pat. No. 6,207,952 (Mar. 27, 2001) describe a water phantom type dose distribution apparatus that includes a closed water tank, filled with water to the brim, having an inserted sensor that is used to determine an actual dose distribution of radiation prior to radiation therapy.

Starting/Stopping Irradiation

K. Hiramoto, et. al. “Charged Particle Beam Apparatus and Method for Operating the Same”, U.S. Pat. No. 6,316,776 (Nov. 13, 2001) describe a charged particle beam apparatus where a charged particle beam is positioned, started, stopped, and repositioned repetitively. Residual particles are used in the accelerator without supplying new particles if sufficient charge is available.

K. Matsuda, et. al. “Method and Apparatus for Controlling Circular Accelerator”, U.S. Pat. No. 6,462,490 (Oct. 8, 2002) describe a control method and apparatus for a circular accelerator for adjusting timing of emitted charged particles. The clock pulse is suspended after delivery of a charged particle stream and is resumed on the basis of state of an object to be irradiated.

Movable Patient

N. Rigney, et. al. “Patient Alignment System with External Measurement and Object Coordination for Radiation Therapy System”, U.S. Pat. No. 7,199,382 (Apr. 3, 2007) describe a patient alignment system for a radiation therapy system that includes multiple external measurement devices that obtain position measurements of movable components of the radiation therapy system. The alignment system uses the external measurements to provide corrective positioning feedback to more precisely register the patient to the radiation beam.

Y. Muramatsu, et. al. “Medical Particle Irradiation Apparatus”, U.S. Pat. No. 7,030,396 (Apr. 18, 2006); Y., Muramatsu, et. al. “Medical Particle Irradiation Apparatus”, U.S. Pat. No. 6,903,356 (Jun. 7, 2005); and Y., Muramatsu, et. al. “Medical Particle Irradiation Apparatus”, U.S. Pat. No. 6,803,591 (Oct. 12, 2004) all describe a medical particle irradiation apparatus having a rotating gantry, an annular frame located within the gantry such that is can rotate relative to the rotating gantry, an anti-correlation mechanism to keep the frame from rotating with the gantry, and a flexible moving floor engaged with the frame is such a manner to move freely with a substantially level bottom while the gantry rotates.

H. Nonaka, et. al. “Rotating Radiation Chamber for Radiation Therapy”, U.S. Pat. No. 5,993,373 (Nov. 30, 1999) describe a horizontal movable floor composed of a series of multiple plates that are connected in a free and flexible manner, where the movable floor is moved in synchrony with rotation of a radiation beam irradiation section.

Respiration

K. Matsuda “Radioactive Beam Irradiation Method and Apparatus Taking Movement of the Irradiation Area Into Consideration”, U.S. Pat. No. 5,538,494 (Jul. 23, 1996) describes a method and apparatus that enables irradiation even in the case of a diseased part changing position due to physical activity, such as breathing and heart beat. Initially, a position change of a diseased body part and physical activity of the patient are measured concurrently and a relationship therebetween is defined as a function. Radiation therapy is performed in accordance to the function.

Patient Positioning

Y. Nagamine, et. al. “Patient Positioning Device and Patient Positioning Method”, U.S. Pat. Nos. 7,212,609 and 7,212,608 (May 1, 2007) describe a patient positioning system that compares a comparison area of a reference X-ray image and a current X-ray image of a current patient location using pattern matching.

D. Miller, et. al. “Modular Patient Support System”, U.S. Pat. No. 7,173,265 (Feb. 6, 2007) describe a radiation treatment system having a patient support system that includes a modularly expandable patient pod and at least one immobilization device, such as a moldable foam cradle.

K. Kato, et. al. “Multi-Leaf Collimator and Medical System Including Accelerator”, U.S. Pat. No. 6,931,100 (Aug. 16, 2005); K. Kato, et. al. “Multi-Leaf Collimator and Medical System Including Accelerator”, U.S. Pat. No. 6,823,045 (Nov. 23, 2004); K. Kato, et. al. “Multi-Leaf Collimator and Medical System Including Accelerator”, U.S. Pat. No. 6,819,743 (Nov. 16, 2004); and K. Kato, et. al. “Multi-Leaf Collimator and Medical System Including Accelerator”, U.S. Pat. No. 6,792,078 (Sep. 14, 2004) all describe a system of leaf plates used to shorten positioning time of a patient for irradiation therapy. Motor driving force is transmitted to a plurality of leaf plates at the same time through a pinion gear. The system also uses upper and lower air cylinders and upper and lower guides to position a patient.

Imaging

P. Adamee, et. al. “Charged Particle Beam Apparatus and Method for Operating the Same”, U.S. Pat. No. 7,274,018 (Sep. 25, 2007) and P. Adamee, et. al. “Charged Particle Beam Apparatus and Method for Operating the Same”, U.S. Pat. No. 7,045,781 (May 16, 2006) describe a charged particle beam apparatus configured for serial and/or parallel imaging of an object.

K. Hiramoto, et. al. “Ion Beam Therapy System and its Couch Positioning System”, U.S. Pat. No. 7,193,227 (Mar. 20, 2007) describe a ion beam therapy system having an X-ray imaging system moving in conjunction with a rotating gantry.

C. Maurer, et. al. “Apparatus and Method for Registration of Images to Physical Space Using a Weighted Combination of Points and Surfaces”, U.S. Pat. No. 6,560,354 (May 6, 2003) described a process of X-ray computed tomography registered to physical measurements taken on the patient's body, where different body parts are given different weights. Weights are used in an iterative registration process to determine a rigid body transformation process, where the transformation function is used to assist surgical or stereotactic procedures.

M. Blair, et. al. “Proton Beam Digital Imaging System”, U.S. Pat. No. 5,825,845 (Oct. 20, 1998) describe a proton beam digital imaging system having an X-ray source that is movable into the treatment beam line that can produce an X-ray beam through a region of the body. By comparison of the relative positions of the center of the beam in the patient orientation image and the isocentre in the master prescription image with respect to selected monuments, the amount and direction of movement of the patient to make the best beam center correspond to the target isocentre is determined.

S. Nishihara, et. al. “Therapeutic Apparatus”, U.S. Pat. No. 5,039,867 (Aug. 13, 1991) describe a method and apparatus for positioning a therapeutic beam in which a first distance is determined on the basis of a first image, a second distance is determined on the basis of a second image, and the patient is moved to a therapy beam irradiation position on the basis of the first and second distances.

Problem

There exists in the art of particle beam treatment of cancerous tumors in the body a need for efficient control of magnet parameters to control acceleration of charged particles in a synchrotron of a charged particle therapy system with reduced power supply requirements. Further, there exists in the art of particle beam therapy of cancerous tumors a need for extraction of charged particles at a specified energy, time, and/or intensity to yield a charged particle beam for efficient, precise, and accurate noninvasive, in-vivo treatment of a solid cancerous tumor with minimization of damage to surrounding healthy tissue in a patient. Still further, there exists a need in the art to continue acceleration of charged particles in a synchrotron during the extraction process.

SUMMARY OF THE INVENTION

The invention comprises a charged particle beam acceleration, extraction, and/or targeting method and apparatus used in conjunction with charged particle beam radiation therapy of cancerous tumors.

DESCRIPTION OF THE FIGURES

FIG. 1 illustrates component connections of a particle beam therapy system;

FIG. 2 illustrates a charged particle therapy system;

FIG. 3 illustrates straight and turning sections of a synchrotron

FIG. 4 illustrates turning magnets of a synchrotron;

FIG. 5 provides a perspective view of a turning magnet;

FIG. 6 illustrates a cross-sectional view of a turning magnet;

FIG. 7 illustrates a cross-sectional view of a turning magnet;

FIG. 8 illustrates magnetic field concentration in a turning magnet;

FIG. 9 illustrates correction coils in a turning magnet;

FIG. 10 illustrates a magnetic turning section of a synchrotron;

FIG. 11 illustrates a magnetic field control system;

FIG. 12 illustrates a charged particle extraction system;

FIG. 13 illustrates two-dimensional scanning of a proton beam focal spot, and

FIG. 14 illustrates three-dimensional scanning of a charged particle beam spot.

DETAILED DESCRIPTION OF THE INVENTION

The invention comprises a charged particle beam acceleration and extraction method and apparatus used in conjunction with charged particle beam radiation therapy of cancerous tumors.

In one embodiment, correction coils are used in combination with magnetic field concentration magnets to stabilize a magnetic field in a synchrotron. For example, high precision magnetic field sensors are used to sense a magnetic field created in one or more turning magnets using winding elements. The sensed magnetic field is sent via a feedback loop to a magnetic field controller that adjusts power supplied to correction coils. The correction coils, operating at a lower power, are capable of rapid adjustment to a new power level. Hence, via the feedback loop, the total magnetic field applied by the turning magnets and correction coils is rapidly adjusted to a new strength, allowing continuous adjustment of the energy of the proton beam. In further combination, a novel extraction system allows the continuously adjustable energy level of the proton beam to be extracted from the synchrotron.

Novel design features of a synchrotron are described. Particularly, turning or bending magnets, edge focusing magnets, magnetic field concentration magnets, winding and correction coils, and extraction elements are described that minimize the overall size of the synchrotron, provide a tightly controlled proton beam, directly reduce the size of required magnetic fields, directly reduce required operating power, and allow continual acceleration of protons in a synchrotron even during a process of extracting protons from the synchrotron.

Cyclotron/Synchrotron

A cyclotron uses a constant magnetic field and a constant-frequency applied electric field. One of the two fields is varied in a synchrocyclotron. Both of these fields are varied in a synchrotron. Thus, a synchrotron is a particular type of cyclic particle accelerator in which a magnetic field is used to turn the particles so they circulate and an electric field is used to accelerate the particles. The synchroton carefully synchronizes the applied fields with the travelling particle beam.

By increasing the fields appropriately as the particles gain energy, the charged particles path can be held constant as they are accelerated. This allows the vacuum container for the particles to be a large thin torus. In practice, it is easier to use some straight sections between the bending magnets and some turning sections giving the torus the shape of a round-cornered polygon. A path of large effective radius is thus constructed using simple straight and curved pipe segments, unlike the disc-shaped chamber of the cyclotron type devices. The shape also allows and requires the use of multiple magnets to bend the particle beam.

The maximum energy that a cyclic accelerator can impart is typically limited by the strength of the magnetic fields and the minimum radius/maximum curvature, of the particle path. In a cyclotron the maximum radius is quite limited as the particles start at the center and spiral outward, thus this entire path must be a self-supporting disc-shaped evacuated chamber. Since the radius is limited, the power of the machine becomes limited by the strength of the magnetic field. In the case of an ordinary electromagnet, the field strength is limited by the saturation of the core because when all magnetic domains are aligned the field may not be further increased to any practical extent. The arrangement of the single pair of magnets also limits the economic size of the device.

Synchrotrons overcome these limitations, using a narrow beam pipe surrounded by much smaller and more tightly focusing magnets. The ability of this device to accelerate particles is limited by the fact that the particles must be charged to be accelerated at all, but charged particles under acceleration emit photons, thereby losing energy. The limiting beam energy is reached when the energy lost to the lateral acceleration required to maintain the beam path in a circle equals the energy added each cycle. More powerful accelerators are built by using large radius paths and by using more numerous and more powerful microwave cavities to accelerate the particle beam between corners. Lighter particles, such as electrons, lose a larger fraction of their energy when turning. Practically speaking, the energy of electron/positron accelerators is limited by this radiation loss, while it does not play a significant role in the dynamics of proton or ion accelerators. The energy of those is limited strictly by the strength of magnets and by the cost.

Charged Particle Beam Therapy

Throughout this document, a charged particle beam therapy system, such as a proton beam, hydrogen ion beam, or carbon ion beam, is described. Herein, the charged particle beam therapy system is described using a proton beam. However, the aspects taught and described in terms of a proton beam are not intended to be limiting to that of a proton beam and are illustrative of a charged particle beam system. Any charged particle beam system is equally applicable to the techniques described herein.

Referring now to FIG. 1, a charged particle beam system 100 is illustrated. The charged particle beam preferably comprises a number of subsystems including any of: a main controller 110; an injection system 120; a synchrotron 130 that typically includes: (1) an accelerator system 132 and (2) an extraction system 134; a targeting/delivery system 140; a patient interface module 150; a display system 160; and/or an imaging system 170.

An exemplary method of use of the charged particle beam system 100 is provided. The main controller 110 controls one or more of the subsystems to accurately and precisely deliver protons to a tumor of a patient. For example, the main controller 110 obtains an image, such as a portion of a body and/or of a tumor, from the imaging system 170. The main controller 110 also obtains position and/or timing information from the patient interface module 150. The main controller 110 then optionally controls the injection system 120 to inject a proton into a synchrotron 130. The synchrotron typically contains at least an accelerator system 132 and an extraction system 134. The main controller preferably controls the proton beam within the accelerator system, such as by controlling speed, trajectory, and timing of the proton beam. The main controller then controls extraction of a proton beam from the accelerator through the extraction system 134. For example, the controller controls timing, energy, and/or intensity of the extracted beam. The controller 110 also preferably controls targeting of the proton beam through the targeting/delivery system 140 to the patient interface module 150. One or more components of the patient interface module 150 are preferably controlled by the main controller 110. Further, display elements of the display system 160 are preferably controlled via the main controller 110. Displays, such as display screens, are typically provided to one or more operators and/or to one or more patients. In one embodiment, the main controller 110 times the delivery of the proton beam from all systems, such that protons are delivered in an optimal therapeutic manner to the patient.

Herein, the main controller 110 refers to a single system controlling the charged particle beam system 100, to a single controller controlling a plurality of subsystems controlling the charged particle beam system 100, or to a plurality of individual controllers controlling one or more sub-systems of the charged particle beam system 100.

Synchrotron

Herein, the term synchrotron is used to refer to a system maintaining the charged particle beam in a circulating path; however, cyclotrons are alternatively used, albeit with their inherent limitations of energy, intensity, and extraction control. Further, the charged particle beam is referred to herein as circulating along a circulating path about a central point of the synchrotron. The circulating path is alternatively referred to as an orbiting path; however, the orbiting path does not refer a perfect circle or ellipse, rather it refers to cycling of the protons around a central point or region.

Referring now to FIG. 2, an illustrative exemplary embodiment of one version of the charged particle beam system 100 is provided. The number, position, and described type of components is illustrative and non-limiting in nature. In the illustrated embodiment, the injection system 120 or ion source or charged particle beam source generates protons. The protons are delivered into a vacuum tube that runs into, through, and out of the synchrotron. The generated protons are delivered along an initial path 262. Focusing magnets 230, such as quadrupole magnets or injection quadrupole magnets, are used to focus the proton beam path. A quadrupole magnet is a focusing magnet. An injector bending magnet 232 bends the proton beam toward the plane of the synchrotron 130. The focused protons having an initial energy are introduced into an injector magnet 240, which is preferably an injection Lamberson magnet. Typically, the initial beam path 262 is along an axis off of, such as above, a circulating plane of the synchrotron 130. The injector bending magnet 232 and injector magnet 240 combine to move the protons into the synchrotron 130. Main bending or turning magnets, dipole magnets, or circulating magnets 250 are used to turn the protons along a circulating beam path 264. A dipole magnet is a bending magnet. The main bending magnets 250 bend the initial beam path 262 into a circulating beam path 264. In this example, the main bending magnets 250 or circulating magnets are represented as four sets of four magnets to maintain the circulating beam path 264 into a stable circulating beam path. However, any number of magnets or sets of magnets are optionally used to move the protons around a single orbit in the circulation process. The protons pass through an accelerator 270. The accelerator accelerates the protons in the circulating beam path 264. As the protons are accelerated, the fields applied by the magnets 250 are increased. Particularly, the speed of the protons achieved by the accelerator 270 are synchronized with magnetic fields of the main bending magnets 250 or circulating magnets to maintain stable circulation of the protons about a central point or region 280 of the synchrotron. At separate points in time the accelerator 270/main bending magnet 250 combination is used to accelerate and/or decelerate the circulating protons while maintaining the protons in the circulating path or orbit. An extraction element of the inflector/deflector system 290 is used in combination with a Lamberson extraction magnet 292 to remove protons from their circulating beam path 264 within the synchrotron 130. One example of a deflector component is a Lamberson magnet. Typically the deflector moves the protons from the circulating plane to an axis off of the circulating plane, such as above the circulating plane. Extracted protons are preferably directed and/or focused using an extraction bending magnet 237 and extraction focusing magnets 235, such as quadrupole magnets along a transport path 268 into the scanning/targeting/delivery system 140. Two components of a scanning system 140 or targeting system typically include a first axis control 142, such as a vertical control, and a second axis control 144, such as a horizontal control. In one embodiment, the first axis control 142 allows for about 100 mm of vertical or y-axis scanning of the proton beam 268 and the second axis control 144 allows for about 700 mm of horizontal or x-axis scanning of the proton beam 268. A nozzle system is optionally used for imaging the proton beam and/or as a vacuum barrier between the low pressure beam path of the synchrotron and the atmosphere. Protons are delivered with control to the patient interface module 150 and to a tumor of a patient. All of the above listed elements are optional and may be used in various permutations and combinations. Use of the above listed elements is further described, infra. Protons are delivered with control to the patient interface module 170 and to a tumor of a patient.

In one example, the charged particle irradiation includes a synchrotron having: a center, straight sections, and turning sections. The charged particle beam path runs about the center, through the straight sections, and through the turning sections, where each of the turning sections comprises a plurality of bending magnets. Preferably, the circulation beam path comprises a length of less than sixty meters, and the number of straight sections equals the number of turning sections. Preferably no quadrupoles are used in or around the circulating path of the synchrotron.

Circulating System

A synchrotron 130 preferably comprises a combination of straight sections 310 and ion beam turning sections 320. Hence, the circulating path of the protons is not circular in a synchrotron, but is rather a polygon with rounded corners.

In one illustrative embodiment, the synchrotron 130, which is also referred to as an accelerator system, has four straight elements and four turning sections. Examples of straight sections 310 include the: inflector 240, accelerator 270, extraction system 290, and deflector 292. Along with the four straight sections are four ion beam turning sections 320, which are also referred to as magnet sections or turning sections. Turning sections are further described, infra.

Referring now to FIG. 3, an exemplary synchrotron is illustrated. In this example, protons delivered along the initial path 262 are inflected into the circulating beam path with the inflector 240 and after acceleration are extracted via a deflector 292 to a beam transport path 268. In this example, the synchrotron 130 comprises four straight sections 310 and four turning sections 320 where each of the four turning sections use one or more magnets to turn the proton beam about ninety degrees. As is further described, infra, the ability to closely space the turning sections and efficiently turn the proton beam results in shorter straight sections. Shorter straight sections allows for a synchrotron design without the use of focusing quadrupoles in the circulating beam path of the synchrotron. The removal of the focusing quadrupoles from the circulating proton beam path results in a more compact design. In this example, the illustrated synchrotron has about a five meter diameter versus eight meter and larger cross-sectional diameters for systems using a quadrupole focusing magnet in the circulating proton beam path.

Referring now to FIG. 4, additional description of the first turning section 320 is provided. Each of the turning sections preferably comprises multiple magnets, such as about 2, 4, 6, 8, 10, or 12 magnets. In this example, four turning magnets 410, 420, 430, 440 in the first turning section 320 are used to illustrate key principles, which are the same regardless of the number of magnets in a turning section 320. A turning magnet 410 is a particular type of circulating magnet 250.

In physics, the Lorentz force is the force on a point charge due to electromagnetic fields. The Lorentz force is given by the equation 1 in terms of magnetic fields with the election field terms not included. F=q(v×B)  eq. 1

In equation 1, F is the force in newtons; B is the magnetic field in Teslas; and v is the instantaneous velocity of the particles in meters per second.

Referring now to FIG. 5, an example of a single magnet turning section 410 is expanded. The turning section includes a gap 510. The gap is preferably a flat gap, allowing for a magnetic field across the gap that is more uniform, even, and intense. The gap 510 runs in a vacuum tube between two magnet halves. The gap is controlled by at least two parameters: (1) the gap 510 is kept as large as possible to minimize loss of protons and (2) the gap 510 is kept as small as possible to minimize magnet sizes and the associated size and power requirements of the magnet power supplies. The flat nature of the gap 510 allows for a compressed and more uniform magnetic field across the gap. One example of a gap dimension is to accommodate a vertical proton beam size of about 2 cm with a horizontal beam size of about 5 to 6 cm.

As described, supra, a larger gap size requires a larger power supply. For instance, if the gap size doubles in vertical size, then the power supply requirements increase by about a factor of four. The flatness of the gap is also important. For example, the flat nature of the gap allows for an increase in energy of the extracted protons from about 250 to about 330 MeV. More particularly, if the gap 510 has an extremely flat surface, then the limits of a magnetic field of an iron magnet are reachable. An exemplary precision of the flat surface of the gap 510 is a polish of less than about five microns and preferably with a polish of about one to three microns. Unevenness in the surface results in imperfections in the applied magnetic field. The polished flat surface spreads unevenness of the applied magnetic field.

Still referring to FIG. 5, the charged particle beam moves through the gap with an instantaneous velocity, v. A first magnetic coil 520 and a second magnetic coil 530 run above and below the gap 510, respectively. Current running through the coils 520, 530 results in a magnetic field, B, running through the single magnet turning section 410. In this example, the magnetic field, B, runs upward, which results in a force, F, pushing the charged particle beam inward toward a central point of the synchrotron, which turns the charged particle beam in an arc. Still referring to FIG. 5, a portion of an optional second magnet turning section 420 is illustrated. The coils 520, 530 typically have return elements or turns at the end of one magnet, such as at the end of the first magnet turning section 410. The return elements take space. The space reduces the percentage of the path about one orbit of the synchrotron that is covered by the turning magnets. This leads to portions of the circulating path where the protons are not turned and/or focused and allows for portions of the circulating path where the proton path defocuses. Thus, the space results in a larger synchrotron. Therefore, the space between magnet turning sections 560 is preferably minimized. The second turning magnet is used to illustrate that the coils 520, 530 optionally run along a plurality of magnets, such as 2, 3, 4, 5, 6, or more magnets. Coils 520, 530 running across turning section magnets allows for two turning section magnets to be spatially positioned closer to each other due to the removal of the steric constraint of the turns, which reduces and/or minimizes the space 560 between two turning section magnets.

Referring now to FIGS. 6 and 7, two illustrative 90 degree rotated cross-sections of single magnet turning sections 410 are presented. The magnet assembly has a first magnet 610 and a second magnet 620. A magnetic field induced by coils, described infra, runs between the first magnet 610 to the second magnet 620 across the gap 510. The gap 510 includes a magnetic field incident surface 670 and a magnetic field exiting surface 680. Return magnetic fields run through a first yoke 612 and second yoke 622. The charged particles run through the vacuum tube in the gap. As illustrated, protons run into FIG. 6 through the gap 510 and the magnetic field, illustrated as a vector, B, applies a force, F, to the protons pushing the protons towards the center of the synchrotron, which is off page to the right in FIG. 6. The magnetic field is created using windings through which a current flows about the core. A first coil makes a first winding coil 650 and a second coil makes up a second winding coil 660. Isolating gaps 630, 640, such as air gaps, isolate the iron based yokes 612, 622 from the gap 510. The gap is approximately flat to yield a uniform magnetic field across the gap, as described supra.

Referring again to FIG. 7, the ends of a single bending or turning magnet are preferably beveled. Nearly perpendicular or right angle edges of a turning magnet 410 are represented by a dashed lines 674, 684. The dashed lines 674, 684 intersect at a point 690 beyond the center of the synchrotron 280. Preferably, the edge of the turning magnet is beveled at angles alpha, α, and beta, β, which are angles formed by a first line 672, 682 going from an edge of the turning magnet 410 and the center 280 and a second line 674, 684 going from the same edge of the turning magnet and the intersecting point 690. The angle alpha is used to describe the effect and the description of angle alpha applies to angle beta, but angle alpha is optionally different from angle beta. The angle alpha provides an edge focusing effect. Beveling the edge of the turning magnet 410 at angle alpha focuses the proton beam.

Multiple turning magnets provide multiple magnet edges that each have edge focusing effects in the synchrotron 310. If only one turning magnet is used, then the beam is only focused once for angle alpha or twice for angle alpha and angle beta. However, by using smaller turning magnets, more turning magnets fit into the turning sections 320 of the synchrotron 310. For example, if four magnets are used in a turning section 320 of the synchrotron, then there are eight possible edge focusing effect surfaces, two edges per magnet. The eight focusing surfaces yield a smaller cross-sectional beam size. This allows the use of a smaller gap 510.

The use of multiple edge focusing effects in the turning magnets results in not only a smaller gap, but also the use of smaller magnets and smaller power supplies. For a synchrotron 310 having four turning sections 320 where each turning sections has four turning magnets and each turning magnet has two focusing edges, a total of thirty-two focusing edges exist for each orbit of the protons in the circulating path of the synchrotron 310. Similarly, if 2, 6, or 8 magnets are used in a given turning section, or if 2, 3, 5, or 6 turning sections are used, then the number of edge focusing surfaces expands or contracts according to equation 2.

$\begin{matrix} {{TFE} = {{NTS}*\frac{M}{NTS}*\frac{FE}{M}}} & {{eq}.\mspace{14mu} 2} \end{matrix}$ where TFE is the number of total focusing edges, NTS is the number of turning section, M is the number of magnets, and FE is the number of focusing edges. Naturally, not all magnets are necessarily beveled and some magnets are optionally beveled on only one edge.

The inventors have determined that multiple smaller magnets have benefits over fewer larger magnets. For example, the use of 16 small magnets yields 32 focusing edges whereas the use of 4 larger magnets yields only 8 focusing edges. The use of a synchrotron having more focusing edges results in a circulating path of the synchrotron built without the use of focusing quadrupoles magnets. All prior art synchrotrons use quadrupoles in the circulating path of the synchrotron. Further, the use of quadrupoles in the circulating path necessitates additional straight sections in the circulating path of the synchrotron. Thus, the use of quadrupoles in the circulating path of a synchrotron results in synchrotrons having larger diameters or larger circumferences.

In various embodiments of the system described herein, the synchrotron has:

-   -   at least 4 and preferably 6, 8, 10, or more edge focusing edges         per 90 degrees of turn of the charged particle beam in a         synchrotron having four turning sections;     -   at least about 16 and preferably about 24, 32, or more edge         focusing edges per orbit of the charged particle beam in the         synchrotron;     -   only 4 turning sections where each of the turning sections         includes at least 4 and preferably 8 edge focusing edges;     -   an equal number of straight sections and turning sections;     -   exactly 4 turning sections;     -   at least 4 edge focusing edges per turning section;     -   no quadrupoles in the circulating path of the synchrotron;     -   a rounded corner rectangular polygon configuration;     -   a circumference of less than 60 meters;     -   a circumference of less than 60 meters and 32 edge focusing         surfaces; and/or     -   any of about 8, 16, 24, or 32 non-quadrupole magnets per         circulating path of the synchrotron, where the non-quadrupole         magnets include edge focusing edges.

Referring again FIG. 6, the incident magnetic field surface 670 of the first magnet 610 is further described. FIG. 6 is not to scale and is illustrative in nature. Local imperfections or unevenness in quality of the finish of the incident surface 670 results in inhomogeneities or imperfections in the magnetic field applied to the gap 510. Preferably, the incident surface 670 is flat, such as to within about a zero to three micron finish polish, or less preferably to about a ten micron finish polish.

Referring now to FIG. 8, additional magnet elements, of the magnet cross-section illustratively represented in FIG. 6, are described. The first magnet 610 preferably contains an initial cross-sectional distance 810 of the iron based core. The contours of the magnetic field are shaped by the magnets 610, 620 and the yokes 612, 622. The iron based core tapers to a second cross-sectional distance 820. The magnetic field in the magnet preferentially stays in the iron based core as opposed to the gaps 630, 640. As the cross-sectional distance decreases from the initial cross-sectional distance 810 to the final cross-sectional distance 820, the magnetic field concentrates. The change in shape of the magnet from the longer distance 810 to the smaller distance 820 acts as an amplifier. The concentration of the magnetic field is illustrated by representing an initial density of magnetic field vectors 830 in the initial cross-section 810 to a concentrated density of magnetic field vectors 840 in the final cross-section 820. The concentration of the magnetic field due to the geometry of the turning magnets results in fewer winding coils 650, 660 being required and also a smaller power supply to the coils being required.

Example I

In one example, the initial cross-section distance 810 is about fifteen centimeters and the final cross-section distance 820 is about ten centimeters. Using the provided numbers, the concentration of the magnetic field is about 15/10 or 1.5 times at the incident surface 670 of the gap 510, though the relationship is not linear. The taper 860 has a slope, such as about 20 to 60 degrees. The concentration of the magnetic field, such as by 1.5 times, leads to a corresponding decrease in power consumption requirements to the magnets.

Referring now to FIG. 9, a turning magnet having magnetic field focusing geometry and correction coils is illustrated.

Focusing Geometry

Referring still to FIG. 9, an additional example of geometry of the magnet used to concentrate the magnetic field is illustrated. As illustrated in FIG. 8, the first magnet 610 preferably contains an initial cross-sectional distance 810 of the iron based core. The contours of the magnetic field are shaped by the magnets 610, 620 and the yokes 612, 622. In this example, the core tapers to a second cross-sectional distance 820 with an angle theta, θ, such as about 10, 20, 30, 45, or 50 degrees. As described, supra, the magnetic field in the magnet preferentially stays in the iron based core as opposed to the gaps 630, 640. As the cross-sectional distance decreases from the initial cross-sectional distance 810 to the final cross-sectional distance 820, the magnetic field concentrates. The smaller angle, theta, results in a greater amplification of the magnetic field in going from the longer distance 810 to the smaller distance 820. The concentration of the magnetic field is illustrated by representing an initial density of magnetic field vectors 830 in the initial cross-section 810 to a concentrated density of magnetic field vectors 840 in the final cross-section 820. The concentration of the magnetic field due to the geometry of the turning magnets results in fewer winding coils 650, 660 being required and also a smaller power supply to the winding coils 650, 660 being required.

Correction Coils

Still referring to FIG. 9, optional correction coils 910, 920 are illustrated that are used to correct the strength of one or more turning magnets. The correction coils 920, 930 supplement the winding coils 650, 660. The correction coils 910, 920 have correction coil power supplies that are separate from winding coil power supplies used with the winding coils 650, 660. The correction coil power supplies typically operate at a fraction of the power required compared to the winding coil power supplies, such as about 1, 2, 3, 5, 7, or 10 percent of the power and more preferably about 1 or 2 percent of the power used with the winding coils 650, 660. The smaller operating power applied to the correction coils 920, 920 allows for more accurate and/or precise control of the correction coils. The correction coils are used to adjust for imperfection in the turning magnets 410, 420, 430, 440.

Referring now to FIG. 10, an example of winding coils and correction coils about a plurality of turning magnets in an ion beam turning section is illustrated. The winding coils preferably cover 1, 2, or 4 turning magnets. In the illustrated example, a winding coil 1030 winds around two turning magnets 410, 420. Correction coils are used to correct the magnetic field strength of one or more turning or bending magnets. In the illustrated example, a first correction coil 1010 corrects a single turning magnet. Combined in the illustration, but separately implemented, a second correction coil 1020 corrects two turning magnets 410, 420. The correction coils supplement the winding coils. The correction coils have correction coil power supplies that are separate from winding coil power supplies used with the winding coils. The correction coil power supplies typically operate at a fraction of the power required compared to the winding coil power supplies, such as about 1, 2, 3, 5, 7, or 10 percent of the power and more preferably about 1 or 2 percent of the power used with the winding coils. The smaller operating power applied to the correction coils allows for more accurate and/or precise control of the correction coils. More particularly, the first correction coil 1010 is used to adjust for imperfection in a magnetic filed produced by the turning magnet 410 or the second correction coil 1020 is used to adjust for imperfection in the turning magnets 410, 420. Optionally, separate correction coils are used for each turning magnet allowing individual tuning of the magnetic field for each turning magnet, which eases quality requirements in the manufacture of each turning magnet.

Correction coils are preferably used in combination with magnetic field concentration magnets to stabilize a magnetic field in a synchrotron. For example, high precision magnetic field sensors 1050 are used to sense a magnetic field created in one or more turning magnets using winding elements. The sensed magnetic field is sent via a feedback loop to a magnetic field controller that adjusts power supplied to correction coils. The correction coils, operating at a lower power, are capable of rapid adjustment to a new power level. Hence, via the feedback loop, the total magnetic field applied by the turning magnets and correction coils is rapidly adjusted to a new strength, allowing continuous adjustment of the energy of the proton beam. In further combination, a novel extraction system allows the continuously adjustable energy level of the proton beam to be extracted from the synchrotron.

For example, one or more high precision magnetic field sensors 1050 are placed into the synchrotron and are used to measure the magnetic field at or near the proton beam path. For example, the magnetic sensors are optionally placed between turning magnets and/or within a turning magnet, such as at or near the gap 510 or at or near the magnet core or yoke. The sensors are part of a feedback system to the correction coils, which is optionally run by the main controller 110. The feedback system is controlled by the main controller 110 or a subunit or sub-function of the main controller 110. Thus, the system preferably stabilizes the magnetic field in the synchrotron elements rather than stabilizing the current applied to the magnets. Stabilization of the magnetic field allows the synchrotron to come to a new energy level quickly.

Optionally, the one or more high precision magnetic field sensors are used to coordinate synchrotron beam energy and timing with patient respiration. Stabilization of the magnetic field allows the synchrotron to come to a new energy level quickly. This allows the system to be controlled to an operator or algorithm selected energy level with each pulse of the synchrotron and/or with each breath of the patient.

The winding and/or correction coils correct 1, 2, 3, or 4 turning magnets, and preferably correct a magnetic field generated by two turning magnets. A winding or correction coil covering multiple magnets reduces space between magnets as fewer winding or correction coil ends are required, which occupy space.

Example II

Referring now to FIG. 11, an example is used to clarify the magnetic field control using a feedback loop 1100 to change delivery times and/or periods of proton pulse delivery. In one case, a respiratory sensor 1110 senses the breathing cycle of the subject. The respiratory sensor sends the information to an algorithm in a magnetic field controller 1120, typically via the patient interface module 150 and/or via the main controller 110 or a subcomponent thereof. The algorithm predicts and/or measures when the subject is at a particular point in the breathing cycle, such as at the bottom of a breath. Magnetic field sensors 1130, such as the high precision magnetic field sensors, are used as input to the magnetic field controller, which controls a magnet power supply 1140 for a given magnetic field 1150, such as within a first turning magnet 410 of a synchrotron 130. The control feedback loop is thus used to dial the synchrotron to a selected energy level and deliver protons with the desired energy at a selected point in time, such as at the bottom of the breath. More particularly, the synchrotron accelerates the protons and the control feedback loop keeps the protons in the circulating path by synchronously adjusting the magnetic field strength of the turning magnets. Intensity of the proton beam is also selectable at this stage. The feedback control to the correction coils allows rapid selection of energy levels of the synchrotron that are tied to the patient's breathing cycle. This system is in stark contrast to a system where the current is stabilized and the synchrotron deliver pulses with a period, such as 10 or 20 cycles second with a fixed period.

The feedback or the magnetic field design coupled with the correction coils allows for the extraction cycle to match the varying respiratory rate of the patient.

Traditional extraction systems do not allow this control as magnets have memories in terms of both magnitude and amplitude of a sine wave. Hence, in a traditional system, in order to change frequency, slow changes in current must be used. However, with the use of the feedback loop using the magnetic field sensors, the frequency and energy level of the synchrotron are rapidly adjustable. Further aiding this process is the use of a novel extraction system that allows for acceleration of the protons during the extraction process, described infra.

Example III

Referring again to FIG. 10, an example of a winding coil 1030 that covers two turning magnets 410, 420 is provided. As described, supra, this system reduces space between turning section allowing more magnetic field to be applied per radian of turn. A first correction coil 1010 is illustrated that is used to correct the magnetic field for the first turning magnet 410. Individual correction coils for each turning magnet are preferred and individual correction coils yield the most precise and/or accurate magnetic field in each turning section. Particularly, the individual correction coil 1010 is used to compensate for imperfections in the individual magnet of a given turning or benign section of the synchrotron 130. Hence, with a series of magnetic field sensors, corresponding magnetic fields are individually adjustable in a series of feedback loops, via a magnetic field monitoring system, which is preferably a subunit of the main controller 110. In one case, an independent correction coil is used for each turning section magnet. Alternatively, a multiple magnet correction coil, such as the second correction coil 1020, is used to correct the magnetic field for a plurality of turning section magnets.

Proton Beam Extraction

Referring now to FIG. 12, an exemplary proton extraction process from the synchrotron 130 is illustrated. For clarity, FIG. 12 removes elements represented in FIG. 2, such as the turning magnets, which allows for greater clarity of presentation of the proton beam path as a function of time. Generally, protons are extracted from the synchrotron 130 by slowing the protons. As described, supra, the protons were initially accelerated in a circulating path 264, which is maintained with a plurality of turning magnets 250. The circulating path is referred to herein as an original central beamline 264. The protons repeatedly cycle around a central point in the synchrotron 280. The proton path traverses through an RF cavity system 1210. To initiate extraction, an RF field is applied across a first blade 1212 and a second blade 1214, in the RF cavity system 1210. The first blade 1212 and second blade 1214 are referred to herein as a first pair of blades.

In the proton extraction process, an RF voltage is applied across the first pair of blades, where the first blade 1212 of the first pair of blades is on one side of the circulating proton beam path 264 and the second blade 1214 of the first pair of blades is on an opposite side of the circulating proton beam path 264. The applied RF field applies energy to the circulating charged-particle beam. The applied RF field alters the orbiting or circulating beam path slightly of the protons from the original central beamline 264 to an altered circulating beam path 265. Upon a second pass of the protons through the RF cavity system, the RF field further moves the protons off of the original proton beamline 264. For example, if the original beamline is considered as a circular path, then the altered beamline is slightly elliptical. The applied RF field is timed to apply outward or inward movement to a given band of protons circulating in the synchrotron accelerator. Each orbit of the protons is slightly more off axis compared to the original circulating beam path 264. Successive passes of the protons through the RF cavity system are forced further and further from the original central beamline 264 by altering the direction and/or intensity of the RF field with each successive pass of the proton beam through the RF field.

The RF voltage is frequency modulated at a frequency about equal to the period of one proton cycling around the synchrotron for one revolution or at a frequency than is an integral multiplier of the period of one proton cycling about the synchrotron. The applied RF frequency modulated voltage excites a betatron oscillation. For example, the oscillation is a sine wave motion of the protons. The process of timing the RF field to a given proton beam within the RF cavity system is repeated thousands of times with successive passes of the protons being moved approximately one micrometer further off of the original central beamline 264. For clarity, the effect of the approximately 1000 changing beam paths with each successive path of a given band of protons through the RF field are illustrated as the altered beam path 265.

With a sufficient sine wave betatron amplitude, the altered circulating beam path 265 touches a material 1230, such as a foil or a sheet of foil. The foil is preferably a lightweight material, such as beryllium, a lithium hydride, a carbon sheet, or a material of low nuclear charge. A material of low nuclear charge is a material composed of atoms consisting essentially of atoms having six or fewer protons. The foil is preferably about 10 to 150 microns thick, is more preferably 30 to 100 microns thick, and is still more preferably 40-60 microns thick. In one example, the foil is beryllium with a thickness of about 50 microns. When the protons traverse through the foil, energy of the protons is lost and the speed of the protons is reduced. Typically, a current is also generated, described infra. Protons moving at a slower speed travel in the synchrotron with a reduced radius of curvature 266 compared to either the original central beamline 264 or the altered circulating path 265. The reduced radius of curvature 266 path is also referred to herein as a path having a smaller diameter of trajectory or a path having protons with reduced energy. The reduced radius of curvature 266 is typically about two millimeters less than a radius of curvature of the last pass of the protons along the altered proton beam path 265.

The thickness of the material 1230 is optionally adjusted to created a change in the radius of curvature, such as about ½, 1, 2, 3, or 4 mm less than the last pass of the protons 265 or original radius of curvature 264. Protons moving with the smaller radius of curvature travel between a second pair of blades. In one case, the second pair of blades is physically distinct and/or are separated from the first pair of blades. In a second case, one of the first pair of blades is also a member of the second pair of blades. For example, the second pair of blades is the second blade 1214 and a third blade 1216 in the RF cavity system 1210. A high voltage DC signal, such as about 1 to 5 kV, is then applied across the second pair of blades, which directs the protons out of the synchrotron through a deflector 292, such as a Lamberson magnet, into a transport path 268.

Control of acceleration of the charged particle beam path in the synchrotron with the accelerator and/or applied fields of the turning magnets in combination with the above described extraction system allows for control of the intensity of the extracted proton beam, where intensity is a proton flux per unit time or the number of protons extracted as a function of time. For example, when a current is measured beyond a threshold, the RF field modulation in the RF cavity system is terminated or reinitiated to establish a subsequent cycle of proton beam extraction. This process is repeated to yield many cycles of proton beam extraction from the synchrotron accelerator.

The benefits of the system include a multi-dimensional scanning system. Particularly, the system allows an energy and/or intensity change while scanning. Because the extraction system does not depend on any change any change in magnetic field properties, it allows the synchrotron to continue to operate in acceleration or deceleration mode during the extraction process. Stated differently, the extraction process does not interfere with synchrotron. In stark contrast, traditional extraction systems introduce a new magnetic field, such as via a hexapole, during the extraction process. More particularly, traditional synchrotrons have a magnet, such as a hexapole magnet, that is off during an acceleration stage. During the extraction phase, the hexapole magnetic field is introduced to the circulating path of the synchrotron. The introduction of the magnetic field necessitates two distinct modes, an acceleration mode and an extraction mode, which are mutually exclusive in time.

Proton Beam Position Control

Referring now to FIG. 13, a beam delivery and tissue volume scanning system is illustrated. Presently, the worldwide radiotherapy community uses a method of dose field forming using a pencil beam scanning system. In stark contrast, FIG. 13 illustrates a spot scanning system or tissue volume scanning system. In the tissue volume scanning system, the proton beam is controlled, in terms of transportation and distribution, using an inexpensive and precise scanning system. The scanning system is an active system, where the beam is focused into a spot focal point of about one-half, one, two, or three millimeters in diameter. The focal point is translated along two axes while simultaneously altering the applied energy of the proton beam, which effectively changes the third dimension of the focal point. For example, in the illustrated system in FIG. 13, the spot is translated up a vertical axis, is moved horizontally, and is then translated down a vertical axis. In this example, current is used to control a vertical scanning system having at least one magnet. The applied current alters the magnetic field of the vertical scanning system to control the vertical deflection of the proton beam. Similarly, a horizontal scanning magnet system controls the horizontal deflection of the proton beam. The degree of transport along each axes is controlled to conform to the tumor cross-section at the given depth. The depth is controlled by changing the energy of the proton beam. For example, the proton beam energy is decreased, so as to define a new penetration depth, and the scanning process is repeated along the horizontal and vertical axes covering a new cross-sectional area of the tumor. Combined, the three axes of control allow scanning or movement of the proton beam focal point over the entire volume of the cancerous tumor. The time at each spot and the direction into the body for each spot is controlled to yield the desired radiation does at each sub-volume of the cancerous volume while distributing energy hitting outside of the tumor.

The focused beam spot volume dimension is preferably tightly controlled to a diameter of about 0.5, 1, or 2 millimeters, but is alternatively several centimeters in diameter. Preferred design controls allow scanning in two directions with: (1) a vertical amplitude of about 100 mm amplitude and frequency up to 200 Hz; and (2) a horizontal amplitude of about 700 mm amplitude and frequency up to 1 Hz. More or less amplitude in each axis is possible by altering the scanning magnet systems.

In FIG. 13, the proton beam goes along a z-axis controlled by the beam energy, the horizontal movement is along an x-axis, and the vertical direction is along a y-axis. The distance the protons move along the z-axis into the tissue, in this example, is controlled by the kinetic energy of the proton. This coordinate system is arbitrary and exemplary. The actual control of the proton beam is controlled in 3-dimensional space using two scanning magnet systems and by controlling the kinetic energy of the proton beam. The use of the extraction system, described supra, allows for different scanning patterns. Particularly, the system allows simultaneous adjustment of the x-, y-, and z-axes in the irradiation of the solid tumor. Stated again, instead of scanning along an x,y-plane and then adjusting energy of the protons, such as with a range modulation wheel, the system allows for moving along the z-axes while simultaneously adjusting the x- and or y-axes. Hence, rather than irradiating slices of the tumor, the tumor is optionally irradiated in three simultaneous dimensions. For example, the tumor is irradiated around an outer edge of the tumor in three dimensions. Then the tumor is irradiated around an outer edge of an internal section of the tumor. This process is repeated until the entire tumor is irradiated. The outer edge irradiation is preferably coupled with simultaneous rotation of the subject, such as about a vertical y-axis. This system allows for maximum efficiency of deposition of protons to the tumor, as defined using the Bragg peak, to the tumor itself with minimal delivery of proton energy to surrounding healthy tissue.

Combined, the system allows for multi-axes control of the charged particle beam system in a small space with low power supply. For example, the system uses multiple magnets where each magnet has at least one edge focusing effect in each turning section of the synchrotron and/or multiple magnets having concentrating magnetic field geometry, as described supra and illustrated in FIG. 13. The multiple edge focusing effects in the circulating beam path of the synchrotron combined with the concentration geometry of the magnets and described extraction system yields a synchrotron having:

-   -   a small circumference system, such as less than about 50 meters;     -   a vertical proton beam size gap of about 2 cm;     -   corresponding reduced power supply requirements associated with         the reduced gap size;     -   an extraction system not requiring a newly introduced magnetic         field;     -   acceleration or deceleration of the protons during extraction;         and     -   control of and/or simultaneous variation of z-axis energy during         extraction.

The result is a 3-dimensional scanning system, x-, y-, and z-axes control, where the z-axes control resides in the synchrotron and where the z-axes energy is variably controlled during the extraction process inside the synchrotron.

Referring now to FIG. 14, an example of a targeting system 140 used to direct the protons to the tumor with 3-dimensional scanning control is provided, where the 3-dimensional scanning control is along the x-, y-, and z-axes. Typically, charged particles traveling along the transport path 268 are directed through a first axis control element 142, such as a vertical control, and a second axis control element 144, such as a horizontal control and into a tumor 1101. As described, supra, the extraction system also allows for simultaneous variation in the z-axis. Thus instead of irradiating a slice of the tumor, as in FIG. 13, all three dimensions defining the targeting spot of the proton delivery in the tumor are simultaneously variable. The simultaneous variation of the proton delivery spot is illustrated in FIG. 14 by the spot delivery path 269. In the illustrated case, the protons are initially directed around an outer edge of the tumor and are then directed around an inner radius of the tumor. Combined with rotation of the subject about a vertical axis, a multi-field illumination process is used where a not yet irradiated portion of the tumor is preferably irradiated at the further distance of the tumor from the proton entry point into the body. This yields the greatest percentage of the proton delivery, as defined by the Bragg peak, into the tumor and minimizes damage to peripheral healthy tissue.

Proton Beam Therapy Synchronization with Breathing

In another embodiment, delivery of a proton beam dosage is synchronized with a breathing pattern of a subject. When a subject, also referred to herein as a patient, is breathing many portions of the body move with each breath. For example, when a subject breathes the lungs move as do relative positions of organs within the body, such as the stomach, kidneys, liver, chest muscles, skin, heart, and lungs. Generally, most or all parts of the torso move with each breath. Indeed, the inventors have recognized that in addition to motion of the torso with each breath, various motion also exists in the head and limbs with each breath. Motion is to be considered in delivery of a proton dose to the body as the protons are preferentially delivered to the tumor and not to surrounding tissue. Motion thus results in an ambiguity in where the tumor resides relative to the beam path. To partially overcome this concern, protons are preferentially delivered at the same point in a breathing cycle.

Initially a rhythmic pattern of breathing of a subject is determined. The cycle is observed or measured. For example, a proton beam operator can observe when a subject is breathing or is between breaths and can time the delivery of the protons to a given period of each breath. Alternatively, the subject is told to inhale, exhale, and/or hold their breath and the protons are delivered during the commanded time period. Preferably, one or more sensors are used to determine the breathing cycle of the individual. For example, a breath monitoring sensor senses air flow by or through the mouth or nose. Another optional sensor is a chest motion sensor attached or affixed to a torso of the subject.

Once the rhythmic pattern of the subject's breathing is determined, a signal is optionally delivered to the subject to more precisely control the breathing frequency. For example, a display screen is placed in front of the subject directing the subject when to hold their breath and when to breath. Typically, a breathing control module uses input from one or more of the breathing sensors. For example, the input is used to determine when the next breath exhale is to complete. At the bottom of the breath, the control module displays a hold breath signal to the subject, such as on a monitor, via an oral signal, digitized and automatically generated voice command, or via a visual control signal. Preferably, a display monitor is positioned in front of the subject and the display monitor displays at least breathing commands to the subject. Typically, the subject is directed to hold their breath for a short period of time, such as about one-half, one, two, or three seconds. The period of time the subject is asked to hold their breath is less than about ten seconds as the period of time the breath is held is synchronized to the delivery time of the proton beam to the tumor, which is about one-half, one, two, or three seconds. While delivery of the protons at the bottom of the breath is preferred, protons are optionally delivered at any point in the breathing cycle, such as upon full inhalation. Delivery at the top of the breath or when the patient is directed to inhale deeply and hold their breath by the breathing control module is optionally performed as at the top of the breath the chest cavity is largest and for some tumors the distance between the tumor and surrounding tissue is maximized or the surrounding tissue is rarefied as a result of the increased volume. Hence, protons hitting surrounding tissue is minimized. Optionally, the display screen tells the subject when they are about to be asked to hold their breath, such as with a 3, 2, 1, second countdown so that the subject is aware of the task they are about to be asked to perform.

A proton delivery control algorithm is used to synchronize delivery of the protons to the tumor within a given period of each breath, such as at the bottom of a breath when the subject is holding their breath. The proton delivery control algorithm is preferably integrated with the breathing control module. Thus, the proton delivery control algorithm knows when the subject is breathing, where in the breath cycle the subject is, and/or when the subject is holding their breath. The proton delivery control algorithm controls when protons are injected and/or inflected into the synchrotron, when an RF signal is applied to induce an oscillation, as described supra, and when a DC voltage is applied to extract protons from the synchrotron, as described supra. Typically, the proton delivery control algorithm initiates proton inflection and subsequent RF induced oscillation before the subject is directed to hold their breath or before the identified period of the breathing cycle selected for a proton delivery time. In this manner, the proton delivery control algorithm can deliver protons at a selected period of the breathing cycle by simultaneously or near simultaneously delivering the high DC voltage to the second pair of plates, described supra, that results in extraction of the protons from the synchrotron and subsequent delivery to the subject at the selected time point. Since the period of acceleration of protons in the synchrotron is constant, the proton delivery control algorithm is used to set an AC RF signal that matches the directed breathing cycle of the subject.

Multi-Field Irradiation

The 3-dimensional scanning system of the proton spot focal point, described supra, is preferably combined with a rotation/raster method. For example, the spot focal point scanning system is optionally used while rotating the individual such as to at least ten distinct irradiation positions. The method includes layer wise tumor irradiation from many directions. During a given irradiation slice, the proton beam energy is continuously changed according to the tissue's density in front of the tumor to result in the beam stopping point, defined by the Bragg peak, to always be inside the tumor and inside the irradiated slice. The novel method allows for irradiation from many directions, referred to herein as multi-field irradiation, to achieve the maximal effective dose at the tumor level while simultaneously significantly reducing possible side-effects on the surrounding healthy tissues in comparison with existing methods. Essentially, the multi-field irradiation system distributes dose-distribution at tissue depths not yet reaching the tumor.

Two additional examples are provided to further describe the cancer therapy system.

In a first example, a particle beam irradiation apparatus for cancer therapy of a tumor of a patient is used for controlling an operational magnetic field, where the operational magnetic field controls direction of charged particles. The apparatus preferably includes a synchrotron having turning magnets, a magnetic field sensor, a gap through which the charged particles pass during use, and at least one correction coil. The turning magnets include windings operating at a first power level to generate a first magnetic field about the gap and at least one correction coil operating at a second power level less than about five percent of the first power level, to generate a correction magnetic field about the gap during use. The first magnetic field and the correction magnetic field combine to yield the operational magnetic field during use. The synchrotron adjusts to a target magnetic field strength via monitoring the operational magnetic field to determine the second power level, where as a result of lower power consumption the second power level allows for more accurate and/or precise control of the operational magnetic field. Preferably, the control of the operational magnetic field is performed using a feedback loop, where the feedback loop uses a signal from the magnetic field sensor to adjust the second power level to obtain the target magnetic field strength.

In a second example, charged particle irradiation of a tumor of a patient is performed by accelerating the charged particle beam with a synchrotron, focusing the charged particles with a focal system to a volume of less than about two millimeters in diameter, controlling kinetic energy of the charged particles using at least one correction coil in said synchrotron; and actively scanning the charged particles at the tumor with an active scanning system. Preferably, the active scanning system simultaneously scans the charged particles along all of: a horizontal axis, a vertical axis, and a kinetic energy axis.

Although the invention has been described herein with reference to certain preferred embodiments, one skilled in the art will readily appreciate that other applications may be substituted for those set forth herein without departing from the spirit and scope of the present invention. Accordingly, the invention should only be limited by the Claims included below. 

1. A particle beam irradiation apparatus for cancer therapy of a tumor of a patient, said apparatus controlling an operational magnetic field, wherein the operational magnetic field controls direction of charged particles, said apparatus comprising: a synchrotron comprising turning magnets, a magnetic field sensor, a gap through which the charged particles pass during use, and at least one correction coil, said magnetic field sensor placed proximate said gap in said synchrotron, wherein, said turning magnets comprise: windings operating at a first power level to generate a first magnetic field about said gap; and said at least one correction coil operating at a second power level, said second power level less than about five percent of said first power level, to generate a correction magnetic field about said gap during use, wherein the first magnetic field and the correction magnetic field combine to yield the operational magnetic field during use, wherein said synchrotron adjusts to a target magnetic field strength via monitoring the operational magnetic field to determine said second power level, wherein as a result of lower power consumption said second power level allows for more accurate and/or precise control of the operational magnetic field.
 2. The apparatus of claim 1, further comprising a feedback loop, said feedback loop using a signal from said magnetic field sensor to adjust said second power level to obtain the target magnetic field strength, wherein the target magnetic field strength is achieved by monitoring the operational magnetic field with said magnetic field sensor, wherein the target magnetic field strength is not achieved by stabilizing current applied to said windings.
 3. The apparatus of claim 1, wherein the target magnetic field strength is continuously adjustable during x-axis and y-axis scanning of the charged particles in targeting the tumor during operation.
 4. The apparatus of claim 3, wherein said second power level controlling said correction coil operates at less than two percent of said first power level applied to said windings of said turning magnets.
 5. The apparatus of claim 1, wherein a single correction coil of said at least one correction coil wraps around at least two of said turning magnets in said synchrotron.
 6. The apparatus of claim 1, wherein said synchrotron further comprises: a focal system focusing the charged particles to a volume of less than about two millimeters in diameter; and an active scanning system, wherein said active scanning system delivers the charged particles to said volume while simultaneously scanning position of said volume along all of: a horizontal axis; a vertical axis; and a kinetic energy axis.
 7. The apparatus of claim 1, wherein said gap comprises an entrance surface through which said operational magnetic field enters, said entrance surface comprising a surface polish of less than about five microns roughness.
 8. The apparatus of claim 7, wherein said gap comprises dimensions of about two centimeters by about five centimeters.
 9. The apparatus of claim 1, wherein at least one of said turning magnets comprises a core, wherein a gap surface cross-section distance of said core comprises a dimension of about five centimeters, wherein a core body cross-section distance of said core running parallel to said gap surface cross-section distance comprises a distance at least one and a half times said surface cross-section distance, wherein the resultant operational magnetic field at said gap surface cross-section distance results from a lesser magnetic field at said core body cross-section distance.
 10. The apparatus of claim 1, wherein at least one of said turning magnets comprises a core wrapped by said at least one correction coil, said gap running through said core, wherein said core expands from an outer dimension of said gap with an angle of about twenty to sixty degrees, wherein the magnetic field concentrates in said core upon approach to said gap.
 11. A method for adjusting an operational magnetic filed of a synchrotron for irradiation of a tumor of a patient with charged particles, wherein the operational magnetic field controls direction of the charged particles, said method comprising the steps of: accelerating the charged particles with a synchrotron, said synchrotron comprising turning magnets, a magnetic field sensor, a gap through which the charged particles pass during use, and at least one correction coil, said magnetic field sensor placed proximate said gap in said synchrotron; operating windings of said turning magnets at a first power level to generate a first magnetic field about said gap; and generating a correction magnetic field with at least one correction coil operating at a second power level, said second power level less than about two percent of said first power level, wherein the first magnetic field and the correction magnetic field combine to yield the operational magnetic field during use; and adjusting the operational magnetic field to a target magnetic field strength via monitoring the operational magnetic field to determine said second power level, wherein as a result of lower power consumption said second power level allows for more accurate and/or precise control of the operational magnetic field.
 12. The method of claim 11, further comprising the step of: monitoring the operational magnetic field with said magnetic field sensor, said magnetic field sensor yielding a magnetic field reading; in a feedback loop, adjusting said second power using said magnetic field reading to obtain the target magnetic field strength.
 13. The method of claim 11, further comprising the step of: continuously adjusting the target magnetic field strength during x-axis and y-axis scanning of the charged particles in targeting the tumor.
 14. The method of claim 11, further comprising the step of: correcting magnetic field strength of at least two of said turning magnets with a single correction coil of said at least one correction coil.
 15. The method of claim 11, further comprising the steps of: accelerating the charged particles to an energy using said synchrotron; focusing the charged particles with a focal system to a volume of less than about two millimeters in diameter; and actively scanning the charged particles at the tumor with an active scanning system, wherein said active scanning system simultaneously scans the charged particles along all of: a horizontal axis; a vertical axis; and a kinetic energy axis.
 16. The method of claim 15, further comprising the step of: concentrating the magnetic field with geometry of a core of at least one of said turning magnets, wherein a gap surface cross-section distance of said core comprises a first distance, wherein a core body cross-section distance of said core running parallel to said gap surface cross-section distance comprises a second distance at least one and a half times said surface cross-section distance.
 17. The method of claim 11, further comprising the step of: concentrating the magnetic field with geometry of a core of at least one of said turning magnets, wherein a correction coil wraps around said at least one of said turning magnets, said gap running through said core, wherein said geometry of said core expands from a magnetic field entrance dimension of said gap with an angle of about twenty to sixty degrees to a second core cross-section distance at least one and a half times larger than said magnetic field entrance dimension of said gap.
 18. A method for irradiation of a tumor of a patient with charged particles, comprising the steps of: accelerating the charged particle beam with a synchrotron; focusing the charged particles with a focal system to a volume of less than about two millimeters in diameter; controlling kinetic energy of the charged particles using at least one correction coil in said synchrotron; and actively scanning the charged particles at the tumor with an active scanning system, wherein said active scanning system simultaneously scans the charged particles along all of: a horizontal axis; a vertical axis; and a kinetic energy axis.
 19. The method of claim 18, further comprising the step of: rotating the patient to at least ten distinct irradiation positions and repeating said steps of accelerating, focusing, controlling, and scanning for each of said irradiation positions. 